The present invention relates to a nuclear medical diagnosis apparatus and in particular, to an image processing in a gamma camera and a single photon emission computed tomography (SPECT) configured by using the gamma camera.
In the nuclear medical diagnosis using a nuclear medical diagnosis apparatus, a medicine marked by a radioactive isotope is applied to an examinee. When the medicine is accumulated in a particular organ or a tumor, a gamma ray is emitted from there. By detecting the gamma ray by a radiation detector (hereinafter, referred to simply as a detector), it is possible to obtain an image based on the medicine distribution.
Moreover, by using medicines of different characteristics, it is possible to perform more accurate nuclear medical diagnosis. For example, by using 99mTc-MIBI capable of obtaining information on the myocardial perfusion and the regional wall motion and 123I-BMIPP capable of obtaining myocardial metabolism information to perform two-nuclide imaging, it is possible to perform more preferable heart function diagnosis.
Moreover, it is advantageous to use one nuclide emitting gamma rays of a plurality of energies such as 201Tl, because it is possible to simultaneously observe two accumulation portions (portions where the administrated medicine is accumulated), i.e., only the vicinity of the body skin by the gamma ray having a low energy and a comparatively deep region by the gamma ray having a high energy. Hereinafter, a multiple nuclide imaging and imaging using a nuclide emitting gamma rays of a plurality of energies will be referred to as “multiple nuclide imaging and the like”.
In the case of the multiple nuclide imaging and the like, contamination of an image corresponding to the gamma ray of a predetermined energy may be caused by a gamma ray of a different energy from the predetermined energy and correction should be made. It should be noted that the contamination means an image corresponding to a gamma ray of other energy than a predetermined energy coming into the image corresponding to the gamma ray of the predetermined energy, which causes degradation of the image corresponding to the predetermined energy. One of the causes which causes the contamination is that the detector does not have an ideal high energy resolution. For example, the gamma ray of 140 keV emitted from 99mTc has energy (photo peak) near to that of the gamma ray of 159 keV emitted from 123I and accordingly, if the energy resolution is insufficient, as shown in FIG. 7, bottoms of peaks 70 and 71 are overlapped on the energy spectrum. To solve this problem, conventionally the effect of the contamination is estimated by quantitatively evaluating, in advance, a contamination component 75 (74) of a total absorption peak 70 (71) of the gamma ray emitted from 99mTc, in a predetermined range (energy window 73) corresponding to the energy of the gamma ray emitted from 123I for example, in case of an imaging only of the image corresponding to the gamma ray of 140 keV emitted from 99mTc (gamma ray emitted from 123I, in a predetermined range (energy window 72) corresponding to the energy of the gamma ray emitted from 99mTc, for example, in case of an imaging only of the image corresponding to the gamma ray of 140 keV emitted from 123I).
Another factor of the contamination is scattering of the gamma ray in an examinee. In the nuclear medical diagnosis apparatus, distribution of the position where the gamma ray is generated is imaged according to the gamma ray coming from directly from the accumulation portion of the administered medicine. On the other hand, when the gamma ray is scattered in the examinee, the advance direction is changed and the information on the position where the administered medicine is accumulated is lost, which causes a noise. Since the gamma ray loses energy by scattering, in the two-nuclide imaging of 99mTc and 123I, for example, if a high-energy (159 keV) gamma ray emitted from the 123I is scattered in the examinee, the gamma ray may cause contamination of the image corresponding to a low-energy (140 keV) gamma ray emitted from the 99mTc. For this contamination, conventionally is used the correction using the technique disclosed in The Journal of Nuclear Medicine, Vol. 34, No. 12, pp. 2216-2221, 1993 and JP-A-7-128450.
On the other hand, recently, study is made on a nuclear medical diagnosis apparatus of pixel type instead of the conventional Anger type. It should be noted that as is shown in FIG. 8, the Anger type detector includes a large-area flat crystal scintillator 81 (such as a monocrystal NaI (TI)) to which an optical device represented by a plenty of photoelectron amplification tubes 82, 83, 84 is attached. On the other hand, the pixel type detector, as shown in FIG. 9, includes small-size prismatic radiation detection devices such as NaI(TI) monocrystal pieces 91, 92, 93, 94, 95 and CdTe semiconductors which are arranged in a rectangular shape. In the case of the NaI(TI) monocrystal pieces, photoelectron amplification tubes 96, 97, 98, 99, 100 are mounted on each of them. In the Anger type detector, a plurality of optical devices (such as photoelectron amplification tubes 82, 83, 84) are mounted on one crystal scintillator 81 and the center of gravity of the light quantity is calculated so as to decide the detection position. However, an error is caused in the detection position by the calculation of the center of gravity. The pixel type is completely different from the Anger type. A signal is read out from each of the pixels (each NaI(TI) monocrystal pieces 91, 92, 93, 94, 95) and accordingly, the detection position decision is stable. Moreover, the pixel type detector using a semiconductor such as CdTe has an excellent energy resolution and attracts a great attention as the nuclear medical diagnosis apparatus of the next generation.
The pixel type detector has a new contamination factor which is not remarkable in the Anger type detector. The factor is scattering of the gamma ray in the detector. As shown in FIG. 10, the gamma ray is not entirely absorbed by the detector 10 but may be scattered. When the scattered gamma ray causes total absorption in another pixel, the total energy is deposited to a plurality of pixels. For example, a gamma ray 21 of 159 keV emitted by 123I causes a scattering 23 in a predetermined pixel and deposits 19 keV while scattered gamma ray 22 of 140 keV is entirely absorbed by other pixel 24. In this case, it is considered that two independent energies of 140 keV and 19 keV are deposited. When the signal of 19 keV is lower than the lower level discrimination (LLD) of a circuit system and cannot be detected, only the signal of 140 keV remains. The output cannot be distinguished from the case when the gamma ray of 140 keV emitted by 99mTc is totally absorbed, which causes contamination of the image corresponding to the gamma ray emitted by the 99mTc.
Even if the radiation source distribution of the gamma ray emitted by the original 123I is point-like, energy is deposited to a pixel other than the target by scattering. This causes contamination of the image corresponding to the gamma ray emitted by the 99mTc. When the administered medicine emits a gamma ray having energy higher than the set of the 123I and 99mTc, the scattered gamma ray is easily transported farther. Accordingly, contamination is generated in the further wider image region. The image corresponding to the gamma ray of the 99mTc by the contamination, i.e., the image erroneously judged to be the image corresponding to the gamma ray emitted by the 99mTc is generated in a pixel different from the radiation source distribution of the gamma ray of the 123I (even though correlation between them is present) and has no relationship with the radiation source distribution of the gamma ray of the 99mTc. This problem cannot be solved by the conventional correction and causes degradation of the image.